Microneedles for transdermal diagnostics: Recent advances and newhorizons
Gui-Shi Liua,1, Yifei Kongb,1, Yensheng Wangb, Yunhan Luoa, Xudong Fanc, Xi Xied,∗,Bo-Ru Yangd,∗∗, Mei X. Wub,∗∗∗
aGuangdong Provincial Key Laboratory of Optical Fiber Sensing and Communications, College of Science & Engineering, Jinan University, Guangzhou, 510632, ChinabWellman Center for Photomedicine, Massachusetts General Hospital, Harvard Medical School, Boston, MA, 02114, USAc Department of Biomedical Engineering, University of Michigan, Ann Arbor, MI, 48109, USAd State Key Laboratory of Optoelectronic Materials and Technologies, School of Electronics and Information Technology, Sun Yat-Sen University, Guangzhou, 510006,China
A R T I C L E I N F O
Keywords:MicroneedlePoint-of-careMinimally invasiveWearable or portable biosensorsHome-based diagnosisContinuous monitoring
A B S T R A C T
Point-of-care testing (POCT), defined as the test performed at or near a patient, has been evolving into acomplement to conventional laboratory diagnosis by continually providing portable, cost-effective, and easy-to-use measurement tools. Among them, microneedle-based POCT devices have gained increasing attention fromresearchers due to the glorious potential for detecting various analytes in a minimally invasive manner. Morerecently, a novel synergism between microneedle and wearable technologies is expanding their detection cap-abilities. Herein, we provide an overview on the progress in microneedle-based transdermal biosensors. It coversall the main aspects of the field, including design philosophy, material selection, and working mechanisms aswell as the utility of the devices. We also discuss lessons from the past, challenges of the present, and visions forthe future on translation of these state-of-the-art technologies from the bench to the bedside.
1. Introduction
Modern medicine has witnessed a continuous growth of needle-based blood collection for laboratory analyses due to high efficiencyand low cost. But underneath the flourish are intractable problems: (1)reuse of un- or inappropriate-sterilized needles is common in devel-oping countries to create a serious risk of transmitting blood-bornepathogens (e.g., HIV, HBV, and HCV) [1]; (2) 3.5%–10% of the world'spopulation expecially children holds a somewhat exaggerated form ofneedle phobia that may cause them to avoid seeking routine andemergency medical care [2]; (3) hypodermic needles and syringes aredifficult for at-home self-administration by untrained personnel withregard to safety and waste management. Even for an easy-to-use bloodglucose meter, frequent finger pricks would result in puncture-relatedpain and discomfort. Prausnitz et al. made remarkable progress in de-veloping an alternative tool by shortening the needle length to around150 μm using microfabrication technology [3]. The revolutionary de-sign of needles, named microneedles (MNs) thereafter, guarantees not
only effective drug transport across the stratum corneum (SC), but alsominimal pain by without hitting nerve endings in the dermis. Sincethen, the research on the development of MNs has moved forward at arapid pace [4–9].
The primary function of MNs is to gain access to biofluids beneaththe skin in a nearly pain-free manner. Human skin is made up of threemain layers: (1) the outermost SC layer having a thickness in the rangeof 10–200 μm [10]; (2) viable epidermis with interstitial fluid (ISF)beneath the SC; (3) dermis composed of blood, lymph vessels, nerveendings, and connective tissue at 300–1500 μm. A typical MN device isinstalled with a single needle or a needle array with a needle length of50–2000 μm, a tip diameter of 1–100 μm, and a base width of25–500 μm [4,6]. The MNs can be fabricated using various materials(e.g., silicon, glass, metals, and polymers) in different structures (solid,hollow, porous, coated, etc.) and shapes (conical, pyramidal, cylind-rical, spiky, snake-fang like, etc.) [6]. So far, the MNs have been mainlyapplied to cosmetics, therapeutics, and diagnostics, among which MNs-based cosmetic treatments travel the furthest distance in practicality to
https://doi.org/10.1016/j.biomaterials.2019.119740Received 28 August 2019; Received in revised form 21 December 2019; Accepted 25 December 2019
be commercialized for over 10 years [8]. At present, MN-based ther-apeutics are in clinical trials and most research are focused on thetransdermal delivery of drugs for influenza vaccination [11–15] or fortreatment of diabetes mellitus [16–18], cancers [19,20], neuropathicpain [21], hair regrowth [22], and obesity [23,24].
It is not until recent years that the MNs have been used as a buildingblock to engineer diagnostic sensors. In general, biosensors for point-of-care testing (POCT) detect medically relevant signals by analyzing ex-ternal secretions (urine, saliva, sweat, and tear) or body fluids (ISF andblood) [25]. The use of external secretions are intrinsically constraineddue to fewer biomarkers at lower concentrations and inaccuracy indisease screening/diagnosis [26,27]. The MNs are deemed ideal bio-sensing platforms since they enable extraction of or access to cutaneousbody fluids in a minimally invasive way for detection of various ana-lytes such as macromolecules, metabolites, and drugs. This new type ofdevice greatly increases the use of ISF as a source of biomarkers forPOCT. Recent studies confirmed that cutaneous ISF has similar profilesof proteins, small molecules, and RNA to blood [28–30]. Additionally,ISF contains some specific biomarkers (e.g., exosomes and residentmemory T cells) that are exclusive or sufficient compared to blood[30,31], which is especially true for cutaneous disorders (e.g., mela-noma, lupus, and psoriasis). Currently, there are two major operatingmodes of MN-based diagnosis: (1) extract ISF or blood that carriesbioanalytes through MNs for follow-up testing; (2) adapt MNs intominuscule capturers or sensors for testing near or at the site of care. Thesecond mode is more advantageous in terms of simplicity and speed,but of narrower applicability at the moment. No matter in which mode,the MNs are rebuilding our expectations of future medical diagnosis.Therefore, it is worthwhile to comprehensively overview the recentadvances in MN-based biosensing research. Herein, the objectives ofthis review are to: (1) summarize the principles of MN-based detection;(2) compare the methods of MN-mediated blood/ISF extraction; (3)expound the utility of MN-based transdermal sensors for bioanalytedetection; (4) discuss the major challenges and future trends in bio-medical applications of MN-based sensors.
2. Diagnosis strategies and principles
We categorize MN-based diagnosis into three modalities based onwhere diagnosis can be made: (1) “Off device”, a MN-device is only fortransdermal sampling of biofluids and the samples are transferred forcentral lab testing; (2) “On device”, a miniaturized analyzer is in-tegrated with a MN-device, so that sample transfer is unnecessary; (3)“On MN”, every single MN mounted in a MN-device functions for in vivobiomarker collector or analyzer. Both “on-device” and “on-MN” modescan be adapted for use in rapid diagnostic tests at the point-of-care.
2.1. Microneedles for transdermal sampling
Biofluids (i.e., ISF and blood) can be drawn out from the skin byMNs based on negative pressure, capillary force, or material absorption.Fig. 1 demonstrates the schematics of hollow (H), porous (P), and solid(S) MNs for skin biofluids sampling [32–34]. Long HMNs of> 1500 μmin length are suitable for blood sampling (see Table 1). They can drawblood from the skin using the capillary force or negative pressure.However, the capillary action is not competent for the situation where alarge volume of blood is required. Therefore, a typical blood samplingdevice comprises a long HMN and an actuator that provides negativepressure to expedite sampling [35]. For short HMNs and PMNs (typicalheight< 900 μm), the capillarity of the continuous microchannelswithin the MN patches is preferable to extract ISF in a self-poweredmanner. In general, the sampling rate based on capillarity can be ex-pedited by reducing the contact angle (θ, see Fig. 1) [36], narrowingthe microchannel diameter (2re) [32], and/or inserting a wick into theHMN [37]. For SMNs, the ISF sampling can be accomplished by fluiddiffusion and material absorption. The SMNs are generally made of
hydrogel that is a polymeric crosslinking network containing in-numerable hydrophilic groups (e.g., –NH2, –OH, and –SO3H) [38]. ISFcan continuously diffuse into the hydrogel SMNs, until the swellingforce and the elastic network retraction force reach equilibrium. Theabsorbed ISF can be unloaded from the MNs by centrifugation and/orsolvent extraction, while the crosslinking network of the hydrogel re-mains intact during manipulations [34,39].
2.2. Microneedles for transdermal detection
The “off device” diagnosis often suffers from insufficient samplingand follow-up time-consuming procedures. In contrast, both “on de-vice” and “on MN” modes avoid the need for sample transfer, whichsignificantly shortens the turnaround time. These two modes have fourcommon types of designs (Fig. 2a–d): (1) attachment of a sensor to thebase of an HMN patch (on device) [40–42]; (2) installment of a sensorinto the lumens of HMNs (on device) [43–48]; (3) modification of SMNsurface to make itself as a sensor (on MN) [49–51]; (4) metallization ofSMNs as dry electrodes (on MN) [52–54].
The working principles for the four designs of MN-biosensors in-clude colorimetry, sandwich immunoassay, enzyme-labeled electro-chemical immunoassay, nucleic acid (NA) recognition, enzymatic/nonenzymic electrochemistry, and skin dry electrodes.
The sensors generally consist of a biorecognition element and anoptical or electrochemical transducer. The enzyme and antibody/an-tigen are mostly used in the MN-based sensors for analytes recognitionand capture. In a colorimetric MN sensor, the enzyme (e.g., glucoseoxidase) is typically used for both biorecognition and catalysis to pro-duce hydrogen peroxide (H2O2) in the presence of analyte. The analyteconcentration is readily estimated from the H2O2-induced color changeof a chromogenic substrate on a paper strip (Fig. 2e) [55]. With respectto the MN immunosensors, the antibody or antigen is generally linkedon the MN surface for in vivo protein capturing. Signal transduction
Fig. 1. Differently shaped and structural MNs for transdermal biofluid samplingdriven by negative pressure, capillarity, or swelling.
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relies on in vitro addition of secondary antibody labeled with either afluorophore for fluorometric assay, or an enzyme for colorimetric assay,such as enzyme-linked immune sorbent assay (ELISA). When transfer-ring ELISA to an electrode to form an electrochemical immunosensor,the analyte concentration can be quantified by monitoring the redoxcurrent between the labeled enzyme and a substrate using the electrode(Fig. 2f). Recently, peptide nucleic acids (PNAs) are immobilized on thehydrogel MNs to specifically bind complementary target DNAs viaWatson-Crick base pairing (Fig. 2g) [56]. The concentration of thePNA/DNA duplex is determined using DNA intercalator either on MN or
off MN after a light-triggered release process. Of note, the MNs com-bined with the colorimetry, ELISA, NA recognition, or electrochemicalimmunoassay are for a single use.
For continuous detection, MNs need to be combined with or mod-ified into electrochemical electrodes which mainly consist of sensingmaterials and conductive electrodes. Most electrochemical MN sensorsare enzyme-based, whereas the other sensors are enzyme-free. Themajority of the enzyme-based MN sensors monitor the formation ofH2O2 during the enzyme-catalyzed reaction of analyte, which causes avariation in the current proportional to the analyte concentration
Table 1Summary of biofluid extraction using different types of MNs.
MN type Materials Length (μm) Sample Sampling volume/rate Sampling time Sampling method Test subject Ref
1500 Blood 31.5 μL – PDMS actuator Rabbit [74]Glass 1500 ISF 20 μL 1–2 h Mechanical pressure Human [30]
PMN Al2O3 900 Water – 15 min Capillary action – [75]PGMA 700 Water – A few seconds Capillary action – [76]
350 ISF – A few seconds Capillary action Human [41]PDMS 600 PBS 1.71 nL min−1 – Compression force – [33]
SMN Metal + paper 750 ISF 2 μL 1 min Paper absorption Wistar rat [77]Metal 1000 Blood 103.5 μL 3 min Vacuum Human [78]
Hydrogel MN PMEV/MA-PEG 600 ISF 0.84 mg 1 h Swelling Rat [79]600 ISF – 5 min Swelling Human [39]600 ISF – 1 h Swelling Porcine [80]
MeHA 800 ISF 1.4 μL 1 min Swelling Mouse [34]Alginate + PLLA 550 ISF 6.5 μL 2 min Swelling – [56]
Fig. 2. Four typical designs of MN-based sensors. (a) HMNs topped with a specific sensor, (b) electrochemical electrode incorporated HMNs, (c) surface-functio-nalized and (d) metalized SMNs. (e–i) Working mechanisms of different MN sensors: colorimetry, immunoassay, nucleic acid recognition, and electrochemistry.
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(Fig. 2h left, so called the first generation of enzymatic biosensor)[25,57]. The H2O2-based sensors are intrinsically influenced by theambient concentration of dissolved oxygen in skin and require highoperating voltage ranging from 0.4 to 0.7 V [44,58–60]. To avoid theseissues, MNs have been combined with the second-generation sensingtechnique [61–63], which utilizes a redox mediator instead of H2O2 toshuttle electrons from the redox center of enzyme to the electrode(Fig. 2h right). For example, a low working voltage of 0.15 V wasachieved for lactate detection using methylene blue as the mediator[62]. While the redox mediator-assisted MN sensors obviate the lim-itation of the first-generation sensing technology, they still suffer fromsusceptibility of enzyme activity to environmental conditions (e.g., pHand temperature). In this context, non-enzymatic catalytic materialshave been introduced to MN electrodes mainly for glucose monitoring[64–66]. The MN sensors utilize the non-enzymatic materials as bothcatalysts and mediators to interact with the analytes and to shuttleelectrons between the catalysts and the electrodes (Fig. 2i). Similardetecting principle applies to the pH and nitric oxide (NO) MN sensors[50,67]. For instance, pH responsive metal oxide can directly transferelectrons from its redox reaction with H3O+ to electrode, which in-duces a potential change for pH detection [68].
Metalized MNs mounting on skin can be used as bioelectrodes forrecording of biopotentials, including electrocardiography (ECG) [69],electroencephalography (EEG) [53], and electromyography (EMG)[70]. The biopotentials are monitored by at least two wet electrodes incontact with the skin; however, the traditional wet electrodes con-taining electrolytes give rise to a high skin impedance and unstablebiopotentials during long-term diagnosis. The impedance of wet elec-trode-skin scheme can be modeled as a combination of an electrolyteresistance (Re) and three impedances of the electrode to electrolyte (Ze),the electrolyte to SC (ZSC), and tissue underneath [71]. The dry MNelectrode can impale the SC to directly access ISF, which circumventsthe need for electrolyte and the high impedance issue arising from theSC, thus effectively eliminating Re, Ze, and ZSC [72]. Compared to thewet electrode, the MN electrode therefore has low impedance to showsuperior sensitivity and long-term stability [52,53,69].
3. Biofluid extraction for off-device diagnosis
3.1. Blood extraction
Blood has long been the most important window through whichclinicians can identify what is happening to a patient's health. Thedermal vasculature is located at 300–1500 μm beneath the skin surface,and therefore the HMN patch with a micro-cannula of> 1500 μm inlength is frequently applied in phlebotomy [9]. Such long HMNs havebeen fabricated using silicon (Si) [81], polymers [82,83], and metals[84,85]. Among the structural materials, metals are mostly used due totheir merits of high Young's moduli and fracture toughness.
A typical HMN-device involves a metal micro-cannula and a va-cuum-generating actuator. Actuators with different operating technol-ogies have been developed for the HMNs, including piezoelectric [86],electrolyte-controlled [36], gel-based [87], thermopneumatic [88], andelastic actuators [89]. The elastic chamber made of poly-dimethylsiloxane (PDMS) is a simple and desirable actuator as negativepressure can be readily induced by elastic deformation. Jung et al. builtup a PDMS chamber with an inlet and an outlet valve [35]. The twovalves allowed the formation of negative pressure, rapid blood extrac-tion, and blood transportation. The team further optimized their designwith a PDMS chamber and a polymer-sealed HMN (Fig. 3a) [73]. TheHMN tip was capped with polyvinyl pyrrolidone (PVP) to form a closedPDMS chamber. With high air permeability of PDMS, negative pressurewas created in the closed chamber after incubation in a vacuum box,followed by depositing an airproofing parylene film to maintain thevacuum. Once the vacuumized device was pressed on skin, the PVPsealer got detached and blood was absorbed into the chamber. Animal
testing confirmed a high rate of blood withdrawal (~7.8 μL s−1). In-spired by the Jung et al.‘s scheme, other groups reported similar HMNmodels without use of sealers or valves [74]. More recently, Blicharzet al. manufactured a blood collection device by enclosing a SMN array,a microchannel system, a vacuum chamber, and a reservoir in a com-pact chassis (Fig. 3b) [78]. The stainless steel SMNs were punched intoand withdrawn from skin by a bi-stable disc spring. As a result of thevacuum actuation, ~100 μL of blood flowed into the reservoir via themicrofluidic channel in 3 min. For the collection process, all a user hasto do is to push the actuation button, so that special training is notmandatory.
3.2. ISF extraction
ISF is formed by blood transcapillary filtration and cleared bylymphatic vessels [90]. It is similar to blood in terms of protein di-versity (93.3% in common) [28], small-molecule metabolite composi-tion (79.3% in common) [29], and RNA profile (92.5% in common)[30]. Therefore, ISF is ideally alternative to blood for clinical diagnosis,along with advantages of good applicability in continuous monitoringdue to its coagulation-free characteristic [91] and high sensitivity to-wards the change of local tissues. In addition, ISF extraction causes lesspain due to the reduction of MN length from 1000–2000 μm to100–800 μm (Table 1). Shorter length provides MNs with higher yieldstrength and shear capacity [92,93], so the short MNs for ISF samplinghave greater flexibility in design and material selection.
3.2.1. Hollow microneedlesHMN holds promise not only in blood extraction, but also in ISF
sampling in a painless manner. Different from extracting blood with asingle needle, MN array is generally employed to withdraw enough ISFavailable in skin (< 1 μL mm−2 [94]). MN patches, with array densityup to 1 × 106 needles cm−2, MN height of 100–400 μm, and innerdiameter of 4–100 μm, have been developed [32,88,95–97]. ShortHMNs for fluid collection have been fabricated using Si [32], polymers[88], or titanium [96], but most of them resort to vacuum to extract ISFand do not demonstrate in vivo sampling. Si, although showing highbrittleness, has been the most common single material for constructionof HMNs [95,98,99], because it allows for accurate microfabrication byphotolithography and endows the HMNs with the self-powered abilitybased on enhanced capillarity. The pioneering work by Mukerjee et al.presented a 20 × 20 array of “volcano-like” HMNs (10 μm in holediameter) on a bulk Si wafer for in vivo sampling ISF [95]. It took15–20 min to transfer ISF from human earlobe to a backside reservoirwith the aid of capillary action. Strambini et al. further narrowed theinner diameter of each HMN to 4 μm and increased the density to1 × 106 needles cm−2 in order to enhance capillary action, so that ISFextraction reached a flow rate of 1 μL s−1 (Fig. 4a) [32]. Recently,cylinder concentric substrate was used to squeeze ISF into a 5-HMNarray via local mechanical pressure [30]. A large amount of ISF (up to20 μL for 1–2 h) from human was successfully collected by the HMN-connected glass capillaries.
The high risk of hole clogging is perceived as a major concern. It hasbeen concluded that Si HMNs with straight side-walls have a higherocclusion probability than those with tapered side-walls [40]. However,the tapered HMNs still have the clogging problem since the pore on theHMN tip tends to cut cells during the insertion. Smith et al. moved thepore to the edge of Si MN shank, forming a snake-fang-like HMN, whichalleviated the plugging issue [95,97]. Another effective method is tocover the micro-channels on a Si strip with a nanoporous membrane(5 μm in thickness) [100].
3.2.2. Porous microneedlesPMNs offer a high-density network of continuous capillary channels
to transfer ISF from the MN-skin interface to the base. Ceramics andmetals were used to manufacture PMNs with hole diameter ranging
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from a few tens of nanometers to several microns [75,101,102]. Theformation of porous ceramic or metal structures generally requires highsintering temperatures (e.g., > 1000 °C). Therefore, porous polymersforming at much lower temperatures are a robust alternative. Porousstructures for constructing MN have been generated by polymerizingacrylic monomers in the presence of porogens [76,103]. For example,by using PEG as a porogen, Liu et al. prepared a PMN from poly(gly-cidyl methacrylate) that was crosslinked by triethylene glycol di-methacrylate and trimethylolpropane trimethacrylate with UV irradia-tion [76]. The PMN patch could trigger biofluid suction by capillaryaction until ISF infiltrated the patch base. Like Si HMNs, the polymerPMNs are also mechanically fragile. Several teams attempted tostrengthen the PMNs through reducing pore diameter and/or density;inevitably, the methods compromised sampling rate and volume[76,101]. More recently, Takeuchi et al. devised an elastic sponge-likePMN using PDMS and hyaluronic acid (HA) to conquer the fragility.The PDMS/HA MN is solid at dry state and turns into an elastic porousstructure after HA dissolving, so that it can absorb fluid by manualcompression and transfer the specimen to a reservoir with the help of acapillary pump (Fig. 4b) [33]. In another case, Prausnitz et al. proposedto insert porous filter papers between adjacent high-strength stainlesssteel SMNs [104]. However, the sandwich-structural MN can merelycollect 3.3 nL ISF in 20 min. Further to this study, the same team re-fined their design by attaching the filter papers onto the backside of theSMN, and thus the collected volume of ISF per min reached> 2 μL[77,105].
3.2.3. Hydrogel microneedlesHydrogel MN is a relatively new type of MNs that is hard under dry
state and swells in skin without degradation. An early hydrogel MNpatch was manufactured through crosslinking poly(methylvinylether/maelic acid) (PMVE/MA) with PEG [106]. The MN was first utilized for
transdermal drug delivery [107], and further used to extract ISF for off-line detection of glucose, lithium, and drugs [39,80]. It required only 5-min insertion to collect sufficient ISF (0.84 μL) from excised porcineskin for analytes detection [79], but the sampling time extended to 1 hfor in vivo experiments [39]. In order to shorten the sampling time,Chang et al. employed a super hydrating HA as a hydrogel backbone[34]. They grafted methacrylate on the HA backbones for crosslinkingand then initiated free radical polymerization with UV radiation to formcrosslinking network (Fig. 4c). The HA MNs successfully reduced theextraction time to 1 min. The large uptake of ISF (1.4 μL) allowed formonitoring subtle changes of ISF glucose/cholesterol levels, showing asimilar trend with the real concentrations in blood. Recently, rapid ISFsampling has also been realized via a hydrogel-coated SMN patch,wherein poly(L-lactide)-made SMNs as a mechanical support pierce theskin and Ca2+-alginate coating swells to extract fluids with a samplingcapacity of ~6.5 μL in 2 min [56,108].
On the whole, significant advances in MN fabrication have de-monstrated good feasibility in biofluid sampling over the past 5 years.Advanced MNs have been able to ultra-rapidly extract blood, such as~100 μL in 3 min using SMNs and> 30 μL in seconds using a singleHMN. As for ISF sampling, using capillary effect or material absorption,the different types of superior MNs have already shorten the extractiontime to 1 min with the ascending order by sampling volume: hydrogelMN (1–3 μL) ≤ PMN (>2 μL)<HMN (> 10 μL) (see Table 1). Thesampling volume of> 1 μL can meet the need of off-line commercialanalysis [34]. Sampling enough biofluids within 1 min already makesMN devices competitive to the traditional needles for blood with-drawal. Among the three kinds of MNs, hydrogel MNs and polymerPMNs are easily fabricated but need additional recovery procedures tounload ISF from the MNs, whereas the ISF extracted by HMNs can beeasily transferred out of the MN lumens to the device backside by ca-pillary action. For this reason, HMN collectors have been further
Fig. 3. Schematics of (a) a capped HMN with a pre-vacuumized PDMS actuator [73] and (b) a SMN array-based device for blood sampling [78]. Figures reprintedwith permissions from Royal Society of Chemistry and Springer Nature.
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developed into on-device sensors of diverse structures, while such de-vices based on hydrogel or porous MNs have been rarely reported. Thefollowing section will describe various integration strategies of HMN-based sensors.
4. On-device diagnosis
4.1. Single-use devices
On-device detection can be realized through affixing a specificanalyzer to an HMN patch. The analyzer is typically either face-to-facemounted on the backside of the HMN patch [32,59], or just placed onthe backside with connecting fluidic channels, through which biofluidsare transferred to the analyzer [89]. Fig. 5a presents a typical HMN-based fluidic platform containing an immuno-modified electrode arrayfor myoglobin/troponin detection [109]. Two syringe pumps enablefluids flow to perform an immunoassay on the antibody-decoratedelectrodes, followed by applying a substrate solution (3,3′5,5′-tetra-methylbenzidine) to the electrodes and a chronoamperometric scan forelectrochemical transduction, giving rise to current responses to bothproteins in the range of 100–1000 ppb. While this kind of MN sensorminimizes an electrochemical immunoassay into a microchip platform,ISF still needs to be transferred to the microelectrodes via the complexmicrofluidic channels. The transfer may be limited by small volume ofaccessible ISF (< 1 μL mm−2 of skin [94]) and increases the detectiontime. To avoid ISF transfer and allow low-volume detection, Ranamu-khaarachchi et al. integrated an immuno-modified metal HMN with anoptofluidic device for vancomycin (VAN) detection (Fig. 5b) [110,111].The lumen of the HMN was modified using the streptavidin-biotin re-action to bind a high density of peptide for VAN recognition. In thisway, the microlumen, as an immunoassay reaction vessel, required only
0.6 nL ISF for analysis. Meanwhile, the high-density peptide allowed alow limit of detection (LoD) of VAN (84 nM) that was estimated fromthe substrate absorbance in the optofluidic channel using a photo-detector.
The HMN-fluidic biosensing platforms show proof-of-concept de-vices for a single use; however, the function modules of pump, bufferand agent reservoir, optical/electronic components involved in theassay are not assembled into the platforms. These components com-plicate the device structure and contribute to the cost of the MN sen-sors, which limits their application as disposable devices. One elegantdesign to avoid the above issues is the assembly of colorimetric strips onthe backside of a HMN patch [89,95,112]. In this design, the fluid isextracted to the testing strip, where analyte is oxidized by enzyme toinduce a color change of substrate (e.g., 3,3′5,5′-tetramethylbenzidine).The concentration can be readily graded by naked eye or quantitativelyevaluated by color analyzer. Due to simplicity of the colorimetric assay,a full-function sensing system can be easily assembled into a compactdevice. Fig. 5c presents such a device system including blood processingand multiple-analyte sensing [89]. It is comprised of a PDMS chamberfor activation, a metal HMN for blood access, and a paper sensor fordiagnostic readout. The PDMS chamber enables the uptake of 30 μLblood onto a sample pad in 30 s. The collected specimen flows througha polysulfone membrane for red blood cell filtration, after which theserum diffuses to two colorimetric reaction zones for glucose andcholesterol. The whole detection procedure is completed within 180 s.While the paper-based MN sensors represent a feasible strategy fordisposable POCT diagnosis, they can only detect small-molecule ana-lytes such as glucose and cholesterol. More studies are needed to exploitthe test papers enabling protein assay for the HMN platform, so as todetect more biomarkers.
Fig. 4. (a) Optical and SEM images of an HMN patch made of a silicon wafer. Protruding length: 100 μm; pitch: 16 μm; external diameter: 9 μm; internal diameter:7 μm; HMN area:> 0.5 × 0.5 cm2 [32]. (b) Schematic of PMNs-based microfluidic chip for ISF extraction and direct analysis. The optical and SEM images show thePMN made of PDMS and hyaluronic acid [33]. (c) Schematic of hydrogel MNs made of a rapidly swelling hyaluronic acid crosslinked by methacrylic anhydride forISF extraction [34]. Figures reprinted with permissions from Elsevier, Springer Nature, and John Wiley and Sons.
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4.2. Continuous monitoring devices
Along with the evolution of disposable MN sensors, a number ofresearch groups have exploited MNs as continuous monitoring plat-forms by incorporating external electrodes. They have developed agamut of device architectures, such as defining serpentine Si nanowireson a SMN tip for pH measurement [113], hitching a nanoelectronicthread electrode to carbon/tungsten MN for long-term electrical re-cording [114], and mounting Ag/AgCl electrodes on a PMN patch foredema monitoring [41].
A commonly used monitoring system is structured by attachingenzymatic electrodes on the base of HMN array [40,99,116]. Fig. 6apresents a typical device for long-term glucose monitoring [59]. Athree-electrode system containing a GOx-coated Pt/C working electrodewas fixed on the backside reservoir of a Si HMN patch. By relieving skinresponse with a citrate-containing buffer preloaded in the reservoir, theprototype worked effectively for a wide concentration range(50–400 mg mL−1) and a long period (up to 72 h), as validated byhuman testing. With similar architecture, Miller et al. extended HMNplatform to K+ sensing by potentiometry, wherein a solid-state ion-selective electrode (ISE) as transducer was connected to the HMN viafluidic microchannels (Fig. 6b) [115]. The ISE was made up of a K+
selective membrane and a 3D porous carbon electrode fabricated byinterferometric lithography. The carbon/K+ ISE responded linearly toK+ (10−5–10−2 M) covering the range of normal physiological levels(3 × 10−3–6 × 10−3 M), indicating that the device held potential inon-body K+ monitoring. If aligned with addressable tailored electrodes,a single patch of HMN array can sense multiple physiological markers.Miller et al. developed a highly multiplexed assay of an HMN array/three tailored electrodes sensing system that could detect glucose, lac-tate, and pH [42].
Moving electrodes into hollow hole represents a modern strategy in
HMN-sensor design. Not only it realizes highly sensitive in situ detec-tion, but also prevents the electrodes from scratch damage during mi-croneedle insertion. An early study reported palladium/aminophenol-modified carbon fibers encased inside the microlumens of a pyramidalHMN array for detecting H2O2 and ascorbic acid [66]. Similarly, Wangand coworkers stacked enzyme-modified SMNs on the top of pyramidalHMNs [58]. The SMNs acting as electrodes were electrodeposited witha poly(o-phenylenediamine) (PPD) layer as an entrapping matrix forglutamate oxidase and GOx, as well as an H2O2 permselective layer toeliminate other electroactive interferences (e.g., ascorbic acid, uric acid,and cysteine) in ISF. This smart design gave rise to a good linear re-sponse to glucose (0–14 mM) or glutamate (0–140 mM). If filled withcarbon paste containing metallic micro/nano-particles and oxidases,the HMNs functioned as transducers [43,117]. The renewable carbonpaste electrodes supported the actual re-usage of HMNs. Moreover, thehigh electrocatalytic activity of metal particles towards H2O2 allowedfor a low operating voltage, e.g., −0.15 V for lactate detection with Rdparticles [43]. Carbon paste was also tailored to construct a self-pow-ered biofuel-cell (BFC) glucose sensor [117]. The sensor consisted of acarbon paste-Pt black cathode for oxygen reduction and a carbon pastebioanode containing tetrathiafulvalene (TTF) mediator and GOx. TheTTF enabled shuttling electrons from the redox reaction at the enzymeto the carbon electrode [118], so that the BFC sensor continuouslyharvested power using glucose as a fuel. The power density of the BFCsensor have a linear response to glucose ranging from 0 to 25 mM.
While filling HMN with electrodes exhibits versatility in workingelectrodes and flexibility in fabrication, these devices do not in-corporate counter electrode and reference electrode into a single HMNpatch, so external electrodes are required to form a two/three-electrodesystem to conduct the electrochemical measurement. Mohan et al.presented an integrated alcohol sensor that assembled a Pt wire-basedworking electrode, a Pt counter electrode, and an Ag/AgCl reference
Fig. 5. (a) HMN chip containing microfluidic channels and electrochemical electrodes [109]. (b) HMN optofluidic sensor for drug monitoring: HMN lumen isimmuno-functionalized to capture vancomycin [110]. (c) HMN combined with a PDMS switch and a paper-based sensor for multi-diagnosis [89]. Figures reprintedwith permissions from John Wiley and Sons and Royal Society of Chemistry.
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electrode into a single MN patch (Fig. 6c) [44]. The Pt wire was de-corated with alcohol oxidase that was sandwiched between size-ex-clusion PPD film and charged-exclusion Nafion film. The dual perm-selective layers eliminated interference of common electroactivemolecules (acetaminophen, uric acid, etc.) and provided a good linearresponse to alcohol from 0 to 80 mM.
5. On-MN diagnosis
5.1. Microneedle immunosensors
Procedures involved in the collection/transfer of an ISF or bloodsample can be omitted through transforming a SMN into an effectivesensor. With the conjugation of antibody/antigen to Si SMNs, Kendallgroup developed a series of Si SMN-based immunosensors for capturingprotein biomarkers from the skin [49,119–124]. They used a hetero-bifunctional carboxyl acid- and sulfhydryl-reactive PEG crosslinker tocreate linkages between capture antibodies/antigens and gold-de-posited Si SMNs. After being applied to the skin for minutes to an hour,the SMNs were peeled off and incubated with dye or enzyme-labeleddetection antibodies. The target biomarkers were then identified andquantified by fluorescence microscope or colorimetric enzyme-sub-strate reactions. This SMN/immunoassay method enabled the sensitivedetection of biomarkers, including IgG antibodies [49,120,124],dengue virus NS1 protein [119], and recombinant P. falciparum(rPfHRP2) [122].
Polymers are biocompatible, inexpensive, and easy-processing ma-terials for preparation of MN immunosensors [125–127]. Althoughpolymer SMNs are easily fabricated using micromolding, the bald MNsurface requires amine functionalization to immobilize antibodies/an-tigens. In one study, Yeow et al. utilized electrophilic aromatic
substitution and NaBH4 reduction reaction to yield an amine-enrichedsurface of polycarbonate (PC) SMNs, to which a heterobifunctional PEGderivative (i.e., NHS-PEG-COOH) was linked for influenza vaccine at-tachment [125]. In another work, antibodies were covalently linked tohexamethylenediamine-modified polylactic acid (PLA) SMNs viahomobifunctional crosslinker glutaraldehyde [126]. The former studydemonstrated the sensitivity comparable to that of the gold-coated SiSMN probe for anti-influenza IgG detection, and the latter achieved alow LoD in proximity to that of ELISA for interleukins (IL)-1α testing.Moreover, the latter MN patch was functionalized in the “One Row-OneAntibody (OPOA)” manner to detect IL-1α and IL-6 simultaneously. Arecent work on photonic crystal (PhC) encoded SMNs took a big stepforward on multiplex detection of biomarkers (Fig. 7) [127]. Differentlycolored PhC balls loaded with a specific type of antibody were en-trapped in different PEGDA-PEG MNs. Thereupon, captured biomarkerswere readily distinguished by reading the refection colors of PhC bar-codes. Their concentrations were determined by measuring the fluor-escence intensity. Simultaneous detection of three inflammatory cyto-kines (TNF-α, IL-1β, and IL-6) was performed on sepsis mice using theSMNs contained three colors of PhC barcodes. The PhC-encoded MNsdemonstrate greater ease than the OPOA [126] and the LEGO-like de-sign—knitting three smaller patches with different antibodies togetherinto a whole patch [122], since the sensing signals from the OPOA andLEGO-like devices provide only the concentration information.
MN-based immunosensing is a promising method to identify proteinmarkers. The capturing of proteins typically requires from 20 min to afew hours to achieve a comparable sensitivity to ELISA [119,121,128].Such long sampling time presents a considerable hurdle to its clinicalapplication. Efforts have been endeavored to enhance sensitivity in ashort time by engineering MN surface and increasing penetrated surfacearea of MN. A first attempt was to minimize non-specific absorption by
Fig. 6. (a) Si HMN patch integrated with three electrochemical electrodes for continuous glucose monitoring [59]. (b) HMN-based microfluidic chip for potassiumions monitoring and SEM images of a 3D porous carbon-based electrode [115]. (c) Pyramidal HMN patch incorporating three electrochemical electrodes for alcoholmonitoring [44]. Figures reprinted with permissions from SAGE, John Wiley and Sons, and Elsevier.
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grafting hetero-bifunctional PEG to gold-coated MNs. The PEG couldresist over 90% of passive absorption of protein [49]. On this basis, theefficiency for capturing target proteins critically relies on the im-mobilization strategy of detecting protein. Protein G was proposed asanchor molecules to align the detecting proteins, which improved thecapture efficiency but reduced surface density of the proteins as com-pared to the EDC/NHS crosslinking. The detecting signals suggestedthat the higher density conquered the better orientation to give a highersensitivity [119]. The sensitivity could be further improved by opti-mizing the EDC/NHS crosslinking for the detecting protein [120]. Inaddition to the surface modifications, Coffey et al. demonstrated thatprolonging MN from 40 μm to 190 μm yielded a 4-fold increase incapture amount of specific IgG, enabling rapid extraction of the bio-marker within 10 min [121]. Shortly thereafter, the same team in-creased the MN density from 20,408 MN cm−2 to 30,000 MN cm−2 andexpanded the array area from 16 mm2 to 32 mm2 in order to enlarge thepenetrated surface area, which obtained a comparable sensitivity in30 s [123].
5.2. Microneedle electrochemical sensors
For non-metal MNs, conductive treatments are generally required toturn MN surfaces into conductive/sensing electrodes for electrical re-cording. Multi-walled carbon nanotube (MWCNT) was early reported toselectively grow on a Si SMN array as a working electrode and a counterelectrode for continuous glucose detection [64]. In fact, most electricalconnections or electrodes were constructed by metal deposition, such asvacuum thermal/electron beam evaporation and sputter coating[53,129–131]. Two types of MNs have been developed to avoid thecumbersome vacuum deposition. One was to blend metal microparticles(palladium) into polymer (PC) solution to directly construct conductiveMNs using a micro-molding process. This strategy provided the com-posite MNs with both catalyzing and monitoring ability towards theferrocyanide redox [132]. The other one was adoption of hydrogel MNas the entrapping matrix of enzyme (GOx or lactose oxidase) andmediator (vinylferrocene). Caliò et al. demonstrated that the swollenhydrogel (PEGDA) MNs as conduits could couple the electrochemicalcurrent of glucose and lactic acid to the electrical contacts on the MNbase [61]. Both strategies greatly simplified the preparation process forthe polymer MN-based sensors, which involved only the micromoldingprocess.
Selective and sensitive detection of biomarkers depends on redoxreactions between target molecules and their enzymes. Enzymes aregenerally immobilized on the conductive SMNs via an electro-poly-merization process, during which the enzymes are entrapped within
conductive (e.g., PPD and PEDOT) [58,60] or non-conductive polymers(e.g., chitosan and polyphenol) [44,129]. These polymers provide astable and biocompatible environment for enzymes, which allows theanalyte of interest accessing to enzymes but restrains escape of the largeenzymes. For instance, GOx within PEDOT matrix remained a goodlinear response to glucose between 2 and 14 mM even after the MNsensor was stored in PBS for 7 days [60]. Recently, for sensitivity en-hancement, various nanomaterials including metal NPs/polymer na-nofibers [133], Cu nanoflowers [134], Au–MWCNTs [62], etc., havebeen applied to MNs as supporting structures of enzymes. The nanos-tructures not only provide large electroactive surface areas, but alsoenhance the conductivities between electrodes and enzymes. For ex-ample, Chen et al. deposited Au/Pt NPs, Pt NPs/polyaniline nanofiber,and dual permselective layers on a stainless steel MN for glucosemonitoring (Fig. 8a) [133], achieving a low LoD (0.1 mM) with a broadlinear range (up to 20 mM), while Bollella et al. construed a sponge-likesurface using MWCNTs and polyMB for lactate detection (Fig. 8b) [62],obtaining a high sensitivity of 1473 μA cm−2 mM with a low LoD of2.4 μM. However, a typical enzymatic MN sensor has to experienceenzyme degradation, complicated immobilization procedure of enzyme,and oxygen limitation.
Non-enzymatic MN sensors have been exploited to overcome thecritical drawbacks of the enzymatic sensors [136]. While Pt catalyst ishighly active for glucose electrooxidation, smooth Pt electrodes aresubject to low sensitivity, poor selectivity, and poisoning by adsorbedintermediates [137,138]. Nanoporous Pt black was electroplated at thetips of stainless steel MNs to enlarge the electroactive surface over 440times larger than that of bare Pt electrode, favoring a high selectivity of1.48 μA mM−1 cm2 towards glucose ranging from 2 to 36 mM [65]. ThePt catalytic surface/ability was further increased by combination of PtNPs and MWCNTs, giving rise to a higher sensitivity of 17.73 μA mM−1
cm2 over the range 3–20 mM [64]. However, biofouling from the for-eign body reaction accumulated on the surface of the Pt electrodesduring a 7-day rabbit test, for which less and less reliable sensing waswitnessed since day 5 [65]. Overcoating a biocompatible membranewould probably improve the biocompatibility of Pt NPs-based elec-trode.
In addition to glucose sensing, electrochemical on-MN sensors haverecently emerged for monitoring NO, pH, and H2O2. NO has been re-cognized as an important messenger molecule under different physio-pathological conditions (e.g., a physiological indicator of colorectalcancer) [139]. Keum et al. mounted a hemin/PEDOT-modified PCLMNs on an endomicroscope to construct an instrument that enabledboth real-time monitoring of NO in colonic polyps and detailed imagingof the lesions [50]. The PCL MN was exquisitely coated with hemin/
Fig. 7. Schematic of (a) photonic crystal encoded MNs and (b) detection of multiple ISF biomarkers [127]. Figure reprinted with permission from John Wiley andSons.
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PEDOT using polydopamine as a binder (Fig. 8c). The hemin offered ap-type doping effect for PEDOT via π-π stacking as well as high specificaffinity toward NO [140]. The presence of NO reduced hemin (Fe3+) toheme (Fe2+), which resulted in current reduction of the PEDOTcoating. Periodical response to NO with micromolar sensitivity wasdemonstrated by alternately applying the MNs in normal and mela-noma tissues. Recent field-effect transistor sensors demonstrated thatconjugation of graphene or reduced graphene oxide (rGO) with por-phyrins (e.g., hemin) significantly promoted selectivity and sensitivityfor NO detection [140,141]. Following this design, Tang et al. synthe-sized iron-porphyrin functionalized graphene (FGPC) and deposited theFGPC on a PEDOT/Au NPs-coated MN [142]. The composite layer en-hanced electron transfer and favored very high specific surface area forNO catalyzing [143,144], endowing the MN sensor with a LoD down tonM level.
Microneedle pH sensors were also developed by surface grafting ofvarious pH-responsive materials [67,130,145,146]. ZnO was reportedas pH sensing layer on a tungsten (W) MN because of its biocompat-ibility, chemical stability, and amphoteric properties [147,148]. TheZnO–W MN sensor could measure pH of 2–9 with a sensitivity of−46.35 mV pH−1, and completed the pH measurement of a mousebrain or bladder within 60 s [67]. A higher sensitivity of −51.2 mVpH−1 was achieved using a composite of molybdenum disulfide (MoS2)nanosheets and polyaniline (PAN) on a stainless steel MN [146]. Theenhanced sensitivity attributed to high specific surface area of MoS2
provided for H+ sensitive PAN. The pH sensitivity could be furtherlifted to a super-Nernst response of 141 mV pH−1 by modifying a W MNwith boron doped diamond (BDD), but the preparation of BBD layerinvolved microwave plasma-assisted chemical vapor deposition [145],leading to a complex fabrication process. Noted that the aforemen-tioned pH responsive materials were constructed on a single metal MNwithout base, having disadvantages of poor reproducibility in devicefabrication and varied penetration depth in measurements. Combina-tion of a MN patch array (MPA) with pH sensing materials could avoidthe issues and, moreover, provide a unique advantage in monitoringspatial distribution of pH, as demonstrated by Zuliani et al. using aniridium oxide (IrOx)-modified MPA for pH distribution mapping of anex vivo rat heart [130].
The MN sensors decorated by nanomaterials exhibit superior sen-sing performances arising from the high specific surface area andelectrocatalytic activity; however, the surface nanostructures on theMNs are susceptible to mechanical damage upon skin insertion. In arecent scheme by Xie et al., a dissolvable polymer PVP was sprayed overMN working electrodes as protective layers (Fig. 8d). The PVP layerwith a thickness of 2 μm could preserve the integrity of vertical ZnOnanowires [135] or nanohybrids of rGO and Pt NPs [149] on the MNsurface during insertion and dissolved in the skin in 5 min. As such, thePVP-protected ZnO/MN sensor was three times more sensitive than thatof the unprotected ZnO/MN sensor for H2O2 monitoring.
Fig. 8. (a) Schematics of (left) functionalized stain-less-steel MN (three main components: Au/Pt NPs,PVDF/Nafion layer, and glucose oxidase) and (right)glucose monitoring [133]. (b) Microscopic structureof (left) pMB-MWCNT composite and (right) SMNsurface (four layers: gold, Au-MWCNTs, poly-methy-lene blue (pMB), and lactate oxidase) for lactate de-tection [62]. (c) Workflow for surface modification ofa polycarprolactone (PCL) MN for real-time NOmonitoring [50]. (d) Schematic of ZnO NWdecoratedstainless steel MNs for H2O2 detection and SEMimages of vertical ZnO NWs synthesized on MN sur-face (top) and protected by PVP (bottom) [135].Figures reprinted with permissions from Elsevier,John Wiley and Sons, and American Chemical So-ciety.
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5.3. Microneedle biopotential recordings
Biopotentials associated with physiological processes are electricalsignals generated inside the body, including, but not limited to, ECG,EEG, EMG, and electrooculogram (EOG). Such biopotentials providebiomarkers for pathological and physiological state and are often re-corded using wet electrodes. However, the conventional wet electrodesfor biopotential recording require tedious procedures, such as skinpreparation and gel usage in order to decrease electrode-skin interfaceimpedance [53,54]. The slow rate of gel drying is prejudicial to long-term monitoring [131,150]. Moreover, motion artifacts arising fromelectrode loosening and skin-potential variation (SPV) distort the sig-nals [69,150]. Thereupon, microneedle electrode arrays (MNEAs), anovel dry electrode, were exploited to address these issues. The MNEAspenetrated through the SC to fasten the electrodes, which providedstable signals in motion states and low impedance density (e.g.,7.5 kΩ cm2@10 Hz) [53]. To further minimize motion artifacts, othertypes of MNEAs were developed on curved [151] or flexible substrates(e.g., parylene [53], SU-8 [52], polyimide [150], and PET [70]) to ac-commodate body movements. Besides, Pei et al. proposed a MN-basedstrategy to eliminate the SPV effect. The main parts of MN electrodeswere overlaid with an insulative parylene layer to avoid contact withthe stratum corneum, while bare conductive MN tips exposed to thestratum germinativum [69]. This method effectively suppressed theSPV interference in a motion state. Furthermore, MN patch with high-density electrode array held a unique advantage in mapping complexmuscle activity with a special pattern [54,151].
6. Conclusions, challenges, and perspectives
Over the past decades, great strides have been made in structural
materials, device designs, and detection strategies to broaden the ap-plication of MN-biosensors. Based on a rich variety of detectionmethods, currently available MN-biosensors are capable of sensing 25analytes (Table 2), and most are small molecules. Suitable target ana-lytes include, but are not limited to glucose, proteins, ions, drugs,metabolites, biopotentials, and plant DNA [152]. With the rapid de-velopment, a few seconds or minutes are enough for biofluid transfer,analyte diffusion, or biorecognition (Table 1). The long collectionperiod for proteins has been reduced to tens of seconds [123]. Recently,hydrogel-coated MNs showed the capablility of transdermal sampling ofNA in 15 min and immune cells in 12 h [56,108]. The NA bound to thehydrogel hybridized with complementary target DNA, the concentra-tion of which was determined by spectrophotometry (Fig. 9a). Asidefrom NA, the hydrogel can be loaded with antigen nanocapsules thatsolicit antigen-presenting cells (APCs) to recruit antigen-specific T cellsinto the hydrogel under the skin (Fig. 9b). With the remarkable pastwork paving the way, the MN-sensor research is on the way to a bril-liant future in spite of many hurdles remaining.
Antibodies and antigens constitute one of the largest and most im-portant classes of disease biomarkers; however, there is relatively a lowequilibrium concentration for most proteins in ISF owing to their lowmicrovascular permeability [153]. Our group found that the illumina-tion on the skin for a few seconds with a 532-nm pulsed laser resulted inthe 1000-fold concentration increase of circulating biomarkers (e.g.,IgG) in the upper dermis [128,154]. When inserting MNs cova-lently linked with influenza antigens into the laser-treated skin, influ-enza antigen-specific IgG was readily detected by the MNs using mouseand pig models receiving influenza vaccines, with the sensitivity, spe-cificity, and accuracy comparable to the immunofluorescence assay ofblood samples [128]. The laser-induced capillary extravasation dependson the preferable absorbance of specific light by hemoglobin, which
Table 2Analytes detected by various MN sensors.
Analyte Sensor structure Detection method Test subject Detection site Ref
Glucose Hydrogel SMNs Glucose assay kit Mouse Off device [34]Metal HMN + Paper sensor Colorimetry Rabbit On device [89]Si HMNs + Sensor Electrochemistry Human On device [40]Metal SMNs + PEDOT Electrochemistry – On MN [60]
Cholesterol Hydrogel SMNs Cholesterol assay kit Mouse Off device [34]Metal HMNs + Paper sensor Colorimetry Rabbit On device [89]
Glutamate Polymer HMNs + SMNs Electrochemistry – On MN [58]Lactate Polymer HMNs + Carbon paste Electrochemistry – On MN [42,43]
Polymer SMNs + MWCNTs Electrochemistry – On MN [62]Influenza IgG Si SMNs + Antigen ELISA Mouse/Swine On MN [49,121,128]NS1 Protein Si MNs + Antibody ELISA Mouse On MN [119]Pf HRP2 Si MNs + Antibody ELISA Mouse On MN [122]IL-1α, IL-6, TNF-α Polymer SMNs + Antibodies ELISA-blotting method Mouse skin, Ex vivo On MN [126]
TroponinPolymer HMNs + Immunoelectrodes Electrochemical immunoassay – On MN [109]
Ascorbic acid Polymer HMNs + Carbon fibers Electrochemistry – On MN [66]K+ Polymer HMNs + ISE Electrochemistry – On device [115]
Metal SMN + ISE Electrochemistry Chicken/porcine skin, Ex vivo On MN [170]NO Polymer SMNs + Hemin + Endomicroscopy Electrochemistry Melanoma mouse On MN [50]
Metal SMN + FGPC Electrochemistry Rat On MN [142]pH Metal SMN + BDD Electrochemistry Mouse stomach, Ex vivo On MN [145]
Metal SMN + ZnO Electrochemistry Mouse On MN [67]Polymer SMN + IrOx Electrochemistry Rat heart, Ex vivo On MN [130]
H2O2 Metal SMN + Pt/rGO Electrochemistry Mouse On MN [ [149]Alcohol Polymer HMNs + Electrodes Electrochemistry Mouse skin, Ex vivo On device [44]Vancomycin Metal HMNs + Peptide ELISA-optofluidic detection – On device [110]β-Lactam Antibiotic Polymer SMNs + β-lactamase Electrochemistry – On MN [51]Organophosphate Polymer HMNs + Carbon paste Electrochemistry Mouse stomach, Ex vivo On device [44, 46]
Levodopa Porcine skin, Ex vivo [48]Tyrosinase [47]DNA-210 Polymer SMNs + PNA oligomers PNA hybridization Human skin, Ex vivo On MN [56]T cell Polymer SMNs + Antigen Nano capsules Immune response Human skin, Ex vivo Off device [108]
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causes the thermal-induced dilation of capillary beneath the skin[155,156]. Likewise, Coffey et al. manually applied the local mechan-ical stimuli (i.e., 1–10 MPa pressure) on the skin to enhance the leakageof proteins through vessels [157]. Increasing the capillary permeabilityseems to be a viable strategy that enables the MNs to measure bloodbiomarkers in the epidermis without damage to blood vessels.
MN-based sensing for POCT is still in an immature phase. Only long-needle-based sensors (≥5 mm in length) for monitoring ISF glucose(e.g., Medtronic Guardian™ Sensor and Abbott's Libre system) have al-ready hit the store shelves [158]. In fact, more MN devices are beingcommercialized as a blood/ISF collector, such as TAP (Fig. 3b, SeventhSense Biosystems Inc.) [78], HemoLink (Tasso Inc.), and an ISF ab-sorber (Renephra Ltd.). Such devices depend on a vacuum to painlesslywithdraw blood/ISF from the skin with MNs. Of note, TAP comprised ofan array of 30 microneedles has received clearance from the U.S. Foodand Drug Administration (FDA) in 2017 and CE Mark in 2018. So far,the FDA has just released a draft guidance to detail when a MN deviceshould be subject to the regulation as a product, but the classificationhas not been made. It may take years or a decade to commercialize MN-based sensors, because it needs to be repeatedly tested for reliability,biosafety, and accessibility prior to clinical use. Design factors affectingthe reliability include the length and surface uniformity of MNs and theapplied mechanical stimulus [123,157]. These factors may also vary
with patients regarding their age, weight, body location, and operativehabit. For instance, finger press led to insufficient penetration for theMNs with ~550 μm in length or penetration failure for those with300 μm in length [159–161]. Although an auxiliary applicator or aforce sensor can ensure successful penetration [162,163], inclusion ofan adequate control to standardize each assay may be more helpful forreliable results. However, whether this can address the reliability issueremains undetermined. On the other hand, both sterilization and dis-posal are in need of particular attention. The risk of infection is highdue to the inappropriate handling and disposal of MNs [164,165]. Forthis reason, manufacturers are required to: (1) attach a disinfectant toMN products; (2) introduce a self-sterilizing coating onto MN surface;or (3) explore an additional enclosure for retracting or blunting MNsafter use [166,167]. For a workable sensor, data collection, processing,and readout modules are indispensable. The design of MN-based bio-sensors that are eligible for commocialization has to take more intoaccount, such as simple fabrication, portable size, affordable price, andadequate reliability.
In our view, research in MN-based sensors will grow rapidly in theupcoming decade. One of the key emerging trends would be the con-vergence of MN-biosensors, optoelectronics, and wireless communica-tion for a comprehensive ubiquitous healthcare solution. For example,Wang et al. built up the first prototype of standalone, wearable MN-
Fig. 9. Schematics of (a) PNA-modified hydrogel/SMN for sampling NA [56] and (b) Alginate-coated SMNs for capturing T cells. Antigen-loaded nanocapsules areencased in the alginate layer [108]. Figures reprinted with permissions from American Chemical Society and The American Association for the Advancement ofScience.
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biosensor comprising HMN arrays (analyte measurement), a flexiblecircuit board (data processing), a Bluetooth module (data transmis-sion), and a battery [47]. The strip-shaped device acts as a bandage toadhere to the skin, which significantly enhances a user's comfort. Si-milar products came up soon on the horizon that requires the MNpatches to perfectly accommodate skin deforming for a good fit[168,169]. It can be envisioned that MN-based biosensors would play acentral role in a comprehensive healthcare service in the future, asdepicted in Fig. 10. The wearable devices should be embedded withvarious smart sensors, therapeutic units, and wireless systems: (1)sensors quantify disease biomarkers and upload patients' medical re-cords to cloud storage; (2) diagnosis is remotely performed by physi-cians or even artificial intelligence; (3) medical treatment decisiontriggers the release of preload drugs directly into the patients' bodies.Apparently, patients would benefit from the decentralization ofhealthcare to manage their own health at home, which may also helpaddress the physician shortage. Overall, it is essential for academia,industry, and government to work closely so as to transform the in-novations in the field into commercial products for better and moreeffective healthcare.
Declaration of competing interest
The authors declare that they have no known competing financialinterests or personal relationships that could have appeared to influ-ence the work reported in this paper.
Acknowledgments
This work was supported by the Henry M. Jackson Foundation(Grant No. 309257-1.00-65282, USA) and Department of Defense/AirForce (FA9550-16-1-00173, USA) to M.X. W, National Natural ScienceFoundation of China (61575084 to Y. L, 61771498 to X. X), Science andTechnology Planning Project of Guangdong Province for Industrial
Applications (2017B090917001, China) and Science and TechnologyProgram of Guangzhou (201803010097, China) to X. X, and theResearch Grants of Sun Yat-Sen University (76120–18821104 to X. X,76120–18843232 to B.R. Y, China).
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